The present invention relates to nuclear magnetic resonance phenomena and in particular to the use thereof in imaging and analysis techniques.
This present application describes the development of a low-cost, robust, and fast, on-line nuclear magnetic resonance (NMR) imager (and associated protocols) suitable for imaging a solid object undergoing continuous translational motion. To date, conventional NMR and MRI measurements on solid objects are performed when they are stationary. This prevents the application of NMR imaging methods to objects moving continuously on conveyor belts, or to semi-solid materials being extruded or otherwise ejected. This severely limits the development of MRI as a sensor in an on-line industrial process. In contrast, the NMR techniques and protocols described in this specification are specifically designed to apply to objects in motion and do not succeed unless the object is translating. This distinguishes the present application from previous NMR and MRI approaches.
Conventional MRI velocity measurements on flowing fluids are a possible exception to the statement that conventional NMR methods are performed only on stationary objects. However the protocols used to image fluid flow are not applicable to solid objects moving with constant velocity. In contrast, the techniques described in the present specification can be applied both to solid translating objects and also to flowing fluids.
An on-line imaging technique which is fast, low-cost, robust and fully automated is important in a number of commercial environments. Some conventional MRI techniques, such as echo planar imaging (EPI), are xe2x80x9cfastxe2x80x9d, with image acquisition times of 100 milliseconds or less, but they require expensive equipment, such as rapidly switched (500 to 2000 Hz), low inductance, strong (10-40 mT mxe2x88x921) gradient generating units, and are not suitable for application in a factory environment and cannot easily be automated. Moreover, motion of an object being imaged during the EPI acquisition time has an adverse effect on EPI image quality. For example, an object of size 5 cm, moving with a velocity of 1 m/s would move its own length (5 cm) if the EPI image acquisition time is 50 ms. In contrast the present specification shows that object motion is essential to the success of the present invention and does not degrade image quality. Moreover the hardware is low cost (relative to today""s commercial NMR spectrometers), robust, and can be fully automated.
The present invention exploits a fundamental physical principle of motional relativity, namely, that a time varying magnetic field (or time-varying field gradient) can be applied to an object in either of two equivalent ways. In the first, conventional, way, the object is stationary and the magnetic field is varied in time. In the second way, exploited by the present invention, the magnetic field (or field gradient) is steady, and instead, the object is moved through the field (or field gradient). The latter way has not, hitherto, been exploited for on-line magnetic resonance imaging.
It is an object of the present invention to provide a method for obtaining magnetic resonance imaging data in respect of an object which is undergoing translational motion.
It is a further object of the present invention to provide apparatus for gathering magnetic resonance imaging data on objects passing therethrough.
It is a further object of the invention to provide a method and apparatus for real time monitoring of objects passing through an imaging unit using magnetic resonance imaging techniques.
According to one aspect, the present invention provides a method of nuclear magnetic resonance imaging comprising the steps of:
conveying an object to be imaged through an imaging module at predetermined velocity, v;
generating, within the imaging module, a spatially characterised, constant magnetic field B0 substantially parallel to the direction of the velocity, v;
generating, within the imaging module, a spatially characterised magnetic field gradient, Gz substantially parallel to the direction of the velocity, v;
generating, within the imaging module, a radiofrequency field B1 pulse transverse to field B0;
detecting nuclear magnetic resonance signals weighted with at least one selected nuclear magnetic resonance parameter from said object.
According to another aspect, the present invention provides an apparatus for gathering nuclear magnetic resonance imaging data comprising:
a first field generating means for generating a spatially characterised, constant magnetic field B0 in an imaging unit volume having a predetermined length along a longitudinal axis thereof, the B0 field being parallel to said longitudinal axis;
a second field generating means for generating, in said imaging unit volume, a spatially characterised magnetic field gradient Gz substantially parallel to B0;
a third field generating means for generating, within the imaging unit volume, radiofrequency field B1 pulses transverse to field B0;
receiver means for detecting nuclear magnetic resonance signals weighted with at least one selected nuclear magnetic resonance parameter from said object;
wherein at least said second field generating means comprises a coil having cylindrical geometry.
According to a further aspect, the present invention provides an apparatus for gathering nuclear magnetic resonance imaging data comprising:
a first field generating means for generating a spatially characterised, constant magnetic field B0 in an imaging unit volume having a predetermined length along a longitudinal axis thereof, the B0 field being parallel to said longitudinal axis;
a second field generating means for generating, in said imaging unit volume, a spatially characterised magnetic field gradient Gz substantially parallel to B0;
a third field generating means for generating, within the imaging unit volume, radiofrequency field B1 pulses transverse to field B0;
receiver means for detecting nuclear magnetic resonance signals weighted with at least one selected nuclear magnetic resonance parameter from said object;
wherein at least said second field generating means comprises a coil having adjacent loops thereof separated by a distance which increases or decreases as a function of the distance along the coil axis.
Embodiments of the present invention will now be described, by way of example, and with reference to the accompanying drawings in which:
FIG. 1 shows a schematic diagram showing principles of a nuclear magnetic resonance imaging apparatus according to the present invention;
FIG. 2 shows a schematic diagram of an exemplary RF field generating unit suitable for use in the present invention;
FIG. 3 shows a schematic diagram of an exemplary Gz field generating unit according to the present invention;
FIG. 4 shows a schematic diagram of an exemplary Gx field generating unit suitable for use in the present invention;
FIG. 5 shows a schematic diagram of an exemplary Gxcfx86 field generating unit suitable for use in the present invention;
FIG. 6 shows an exemplary pulse sequence suitable for T2 weighting based on motionally modified spin echoes;
FIG. 7 shows an exemplary pulse sequence suitable for T1 weighting based on motionally modified inversion recovery;
FIG. 8 shows an exemplary pulse sequence suitable for T1 and diffusive weighted imaging based on motionally modified stimulated echoes;
FIG. 9 shows an exemplary pulse sequence suitable for diffusion weighting based on motionally modified spin echoes;
FIG. 10 shows an exemplary pulse sequence suitable for T1 and diffusive weighted imaging based on motionally modified stimulated echoes;
FIG. 11 shows an exemplary arrangement suitable for weighting motional echoes with T1 (low field) relaxation;
FIG. 12 shows an exemplary pulse sequence suitable for three-dimensional imaging based on motional echoes;
FIG. 13 shows an on-line variation of an echo planar imaging pulse sequence of imaging in the x-y plane;
FIG. 14 is a plot of signal intensity versus time showing pseudo-echoes as generated by a computer simulation of free induction decays which are not produced in real measurements;
FIGS. 15 and 16 show deconvolved image profiles as a result of processing the pseudo-echoes of FIG. 14, with FIG. 15 corresponding to the first pseudo-echo and FIG. 16 corresponding to the second pseudo-echo;
FIGS. 17 to 24 show the results of converting a motionally modified free induction decay signal into an image projection in which FIG. 17 shows the signal intensity values of the original profile;
FIG. 18 shows the signal intensity values of a transformed echo with no phase error;
FIG. 19 shows the signal intensity values of a transformed echo with phase error;
FIG. 20 shows the signal intensity values of a transformed second half of an echo with no phase error;
FIG. 21 shows the sign intensity values of a transformed second half of an echo with phase error;
FIG. 22 shows the signal intensity values of a transformed symmetrized echo with no phase error;
FIG. 23 shows the signal intensity values of a transformed symmetrized echo with phase error;
FIG. 24 shows the signal intensity values of a transformed symmetrized echo with phase correction;
FIG. 25 shows a numerical evaluation of the real component of the motional phase factor exp{ixcex3G.vt2/2} as a function of acquisition time;
FIG. 26 shows a simulated signal of a moving object including an initial motionally modified free induction decay and the first two motionally modified spin echoes;
FIG. 27 shows the signal intensity values of a deconvolved noiseless image profile from the free induction decays of FIG. 26;
FIG. 28 shows the signal intensity values of a deconvolved noiseless image profile from the first spin echo of FIG. 26;
FIG. 29 shows the signal intensity values of a deconvolved noiseless image profile from the second spin echo of FIG. 26;
FIG. 30 shows a simulation of the magnetic field direction as a function of x and z spatial co-ordinates for a Gz unit of the present invention;
FIG. 31 shows a simulation of the magnetic field strength Bz along the z axis as a function of z, with x=0, for the Gz unit of FIG. 30;
FIG. 32 shows a simulation of the magnetic field strength Bz as a function of x, with z=0, for the Gz unit of FIG. 30;
FIG. 33 shows a simulation of the magnetic field strength Bz as a function of x, with z=0, for the Gz unit of FIG. 30;
FIGS. 34 and 35 show a simulation of the magnetic field strength B2 as a function of x, with z=xe2x88x92L/2, for the Gz unit of FIG. 30;
FIGS. 36 and 37 show a simulation of die magnetic field strength Bz as a function of x, with z=+L/2, for the Gz unit of FIG. 30;
FIG. 38 shows a simulation of the magnetic field strength Bx perpendicular to the z axis as a function of x, with z=0, for the Gz unit of FIG. 30;
FIG. 39 shows a simulation of the magnetic field strength Bz as a function of x, with z=0, for the Gz unit of FIG. 30;
FIG. 40 shows a simulation of the magnetic field direction and strength as a function of x and z spatial co-ordinates for a Gx unit of the present invention;
FIG. 41 shows a simulation of the magnetic field strength Bz along the axis of a current loop as a function of z, at x=0, for the Gx unit of FIG. 40;
FIG. 42 shows a simulation of the magnetic field strength Bz in the plane of a current loop as a function of x, at z=0, for the Gx unit of FIG. 40;
FIG. 43 shows a simulation of the magnetic field strength Bz in the plane of a current loop as a function of x, at z=0, for the Gx unit of FIG. 40;
FIG. 44 shows a simulation of the square root of magnetic field strength Bzxe2x88x921 in the plane of a current loop as a function of x, at z=0, for the Gx unit of FIG. 40;
FIG. 45 shows a schematic diagram illustrating distance travelled by an object during excitation and acquisition using a Hahn echo pulse sequence;
FIG. 46 shows the distortion in the output as a function of z derived from transforming the first motionally modified spin echo in a Hahn spin echo sequence, from a rectangular object, where the gradient field deviates from the ideal value by +1%;
FIG. 47 shows the distortion in the output as a function of z derived from transforming the first motionally modified spin echo in a Hahn spin echo sequence, from a rectangular object, where the gradient field deviates from the ideal value by xe2x88x922%; and
FIG. 48 shows an undistorted output as a function of z derived from transforming the first motionally modified spin echo in a Hahn spin echo sequence, from a rectangular object where the gradient field is at the ideal value.
a) Motionally Modified Free Induction Decays (MMFID""s) and Motionally Modified Spin-echoes (MMSE""s)
The present invention exploits what can be called xe2x80x9cmotionally modified free induction decaysxe2x80x9d (MMFID""s), and motionally modified spin echoes (MMSE""s). We therefore begin with a description of how an MMFID and an MMSE are formed and how they can be exploited for on-line imaging.
Consider a solid object moving in a straight line with uniform velocity v. Preferably, the first step in the on-line method is to induce longitudinal magnetisation Mz in the object in the same direction as the velocity vector v. The second step involves acquiring the Free Induction Decay (FID) by irradiating with a hard 90xc2x0, on-resonance, radiofrequency pulse in a static, homogeneous magnetic field B0, oriented parallel to the velocity vector v, and in the presence of a linear magnetic field gradient Gz also oriented parallel to B0 and v. Then the FID will be modulated by the motion and attenuated by transverse relaxation T2*.
The existence of MMFID""s can be demonstrated both using analytic mathematical methods (as described in greater detail in Appendix 1) and by computer simulation (as described in greater detail in Appendix 2, where a novel computer algorithm is presented for extracting an image projection from the motionally-modified FID).
The existence of MMSE""s can also be demonstrated using analytic mathematical methods (as described in greater detail in Appendix 3) and by computer simulation (as described in greater detail in Appendix 4). An MMSE is created with the spin echo pulse sequence 90-xcfx84-180-xcfx84-MMSE, when the object to be imaged is moving with constant velocity v in the presence of a constant, linear field gradient Gz oriented parallel to v. In Appendix 3 it is shown that, contrary to the conventional spin-echo sequence, there is no spin echo for an arbitrary pulse spacing xcfx84 because of destructive dephasing by motion through the field gradient. However, provided the pulse spacing xcfx84 is set equal to a multiple of the acquisition time AQ a motionally modified spin echo (MMSE) can be formed. Setting xcfx84 equal to AQ is a necessary, but not a sufficient condition for formation of MMSE""s. The dwell time and gradient also need matching with a sample velocity and sample length. The method for doing this is explained in Appendix 8. Clearly, a train of MMSE""s can be created using successive 180 pulses in the sequence 90-xcfx84-(180-xcfx84-MMSE)n where nxe2x89xa71 and xcfx84 is a multiple of AQ. Appendix 4 presents a novel computer algorithm for extracting an image from an MMSE.
b) Hardware Requirements for Creating MMFID""s and MMSE""s
The objects to be imaged travel in a single-file manner down a conveyor tube, pneumatic tube, belt or other suitable means, schematically shown on FIG. 1 as conveyor 1. The conveyor 1 and all the objects 2 on it preferably move with a constant velocity v. Although the imaging procedure is tolerant of small vibrations of each object (see below), there should be substantially no reorientation of the objects 2 relative to the conveyor 1. In practice this can be arranged by simply holding the objects, in e.g. foam supports (not shown), along the conveyor 1.
The hardware required to create and observe motional echoes consists of separate cylindrically shaped units which enclose the conveyor and can be positioned at various positions along the conveyor. The conveyor 1 carrying the objects to be imaged then travels down the central axis of the cylindrical units, although precise lateral positioning of the objects within the cylindrical units is not essential where uniform fields across the x and y axes are used. A modular approach to the design of the hardware provides greater adaptability to a plurality of applications.
Because the object to be imaged is travelling with velocity v and it takes a finite time (at least AQ) to acquire the NMR signal(s), it is necessary that the B0 field, radiofrequency field B1, and gradient field Gz, are all spatially homogeneous over a distance of, at least vxc3x97AQ, along the conveyor.
For the Hahn echo sequence consisting of a 90 degree radiofrequency excitation pulse followed by a 180 degree pulse a time 2xc3x97AQ later, the distance travelled by the object between excitation and echo acquisition is actually 3vxc3x97AQ. This is illustrated in FIG. 45. This means that B0 field, the radiofrequency field, B1, and the gradient field, Gz, are all preferably homogeneous over a distance of at least 4vxc3x97AQ in order to encompass the whole object during its motion.
According to the preferred embodiment illustrated, the hardware devices as described below for generating the B0 field, radiofrequency field B1, and field gradient Gz is that they are all cylindrically shaped with lengths that can be extended indefinitely, at least in principle. This distinguishes them from conventional NMR arrangements, such as U-shaped magnets, radiofrequency Helmholtz coils, birdcage coils etc. which would create homogeneous fields only over a limited distance along the conveyor.
The Polarizer Unit
We consider a solid object moving with constant velocity v. The first step in obtaining an image of the object is to induce longitudinal magnetisation within it by application of a constant external magnetic field. This is done in the polarizer unit 3 illustrated in FIG. 1. If the object has a short T1, then the polarizer consists either of a single, straight, hollow, cylindrical permanent magnet 3a of length L, as shown in the FIG. 1 inset, or a solenoid coil electromagnet of length L. The object moves on the conveyor 1, preferably down the central axis inside the cylindrical polarizer. The time spent inside the polarizer is L/v and for 100% polarization it is preferred that this should be at least 5T1. However, 100% polarization is not an essential requirement of the on-line imager and lower degrees of polarization can be used.
If T1 is long (several seconds) L may be impracticably large, in which case a series of solenoids or permanent magnets can be arranged in, for example a spiral arrangement and the conveyor passed along the spiral. Once the object is sufficiently polarized it passes, with velocity v, into the imaging module which creates MMFID""s.
The Imaging Module
Depending on the application, an imaging module 4 consists of some or all of five different hardware units. These are called the B0 unit; the RF unit; the Gz unit; the Gx unit; and the Gxcfx86 unit respectively. To create MMFID""s or MMSE""s, only the three B0, RF and Gz units are required.
The B0 Unit
The polarized object 2 emerging from the polarizer 3 on the conveyor 1 passes into a spatially uniform, constant magnetic field B0 created by the B0 unit within the imaging unit 4. Like the polarizer unit, the B0 field can be created by a hollow cylindrical permanent magnet or by a hollow cylindrical solenoid electromagnet called, for convenience, the B0 unit. The conveyor carrying the polarized objects then moves down the middle of the cylinder with uniform velocity in a direction parallel to the cylinder axis and preferably along the cylinder axis. The direction of the polarized longitudinal magnetisation in the object leaving the polarizer should be in the same direction as B0 in the B0 unit. The magnet can be of any desired length provided the B0 field everywhere in the object is spatially uniform. if the object T1 is sufficiently short then the polarizer and B0 units can be combined into a single continuous unit.
The RF Unit
This is illustrated in FIG. 2. The on-resonance, radiofrequency field B1, which must be transverse to B0, can be generated by the specially designed, cylindrical xe2x80x9cradiofrequency solenoid-like coilxe2x80x9d which we call the xe2x80x9cRF unitxe2x80x9d. Preferably this also acts as a receiver coil and its particular preferred form is fully described in the reference, xe2x80x9cA solenoid-like coil producing transverse radiofrequency fields for MR imagingxe2x80x9dby E. K. Jeong, D. H. Kim, M. J. Kim, S. H. Lee, J. S. Suh and Y. K. Kwong in J. Magn. Reson. 127 (1997) 73-79. Article no. MN971172.
The RF unit 20 as described therein includes a pair of cylindrical coils: a first, outer coil 21 which has the plane of each loop 22 tilted with respect to the cylinder (z) axis to generate an RF field with a component perpendicular to the cylinder axis. A second, inner coil 23 acts as an eddy-current coil which eliminates the longitudinal component of the RF field, leaving an RF field entirely perpendicular to the cylinder axis.
A special characteristic of this device is the generation of a uniform radiofrequency field over a long z distance. This distinguishes it from more conventional RF generators such as the standard saddle coil, birdcage or cavity resonator. Such conventional devices could be used for the purposes of the present invention, provided they are of sufficient size that their RF field is uniform over distances exceeding the distance moved by the object during the acquisition time (3vxc3x97AQ, in a spin echo imaging experiment, where v is the velocity and AQ is the acquisition time, see FIG. 45). The solenoid-like RF unit coil 20a overcomes this limitation and can be easily situated inside and concentric with the B0 unit solenoid coil or permanent magnet. The RF unit 20 is interfaced with conventional electronic equipment and computers for control, acquisition and image processing.
The Gz Unit
This is illustrated in FIG. 3. The linear magnetic field gradient Gz is oriented parallel to B0 and the direction of object motion and is generated by the Gz unit. Preferably, this comprises a specially designed, non-uniformly wound cylindrical solenoid coil 30 as shown, in which the spacing of the turns 31 of the coil 30 vary as a function of z-position. This is described more fully in Appendix 5 which includes computer calculations of the field gradient G within the unit. Note that the gradient will need to be matched to the sample velocity and other parameters, so that a suitably sized image is created. The factors determining the magnitude of the Gz gradient magnitudes are detailed in Appendix 8. It may also be possible to create the extended linear Gz gradient using permanent magnets.
The gradient solenoid unit 30 will be located inside and concentrically with the B0 unit and around the radiofrequency solenoid-like RF unit. It can be made of any desired length, subject to a minimum length below which the gradient is no longer sufficiently uniform for undistorted image acquisition.
The Gx Unit
This is illustrated schematically in FIG. 4 and the magnetic fields are simulated in Appendix 6. Where diffusion-weighted on-line imaging is required, an optional fourth type of hardware unit comprising a single electric coil 40 can be used which surrounds the B0 unit 41. The Gx unit creates a steady, spatially localised, non-uniform magnetic field gradient Gx transverse to B0. The coil 40 is wound around the gradient (Gz unit 30) or RF solenoid unit 20 at a single location.
The Gxcfx86 Unit
With reference to FIG. 5, where 2- or 3-dimensional imaging is required (as apart from one-dimensional projection imaging along the direction of motion, z), additional linear, steady, magnetic field gradients transverse to B0 can be created in a conventional way with, for example, carrying imbalanced currents coils 50, 51 placed at suitable locations around the B0 unit 52 magnet. The optional additional coils needed to create these gradients, we call a Gxcfx86 unit because the field can be oriented at an angle xcfx86 to the vertical.
Triggering the NMR Pulse Sequence
This can be done electronically by arranging for the object to be imaged to cut a laser or infra-red beam traversing the conveyor as the object enters the imaging unit. This is provided by suitable conventional light source 5 and a receiver unit 6 coupled to a control circuit 7. An electronic delay then triggers the first 90xc2x0 radiofrequency pulse. If two laser beams are used spaced along the conveyor, the velocity v of the object can be measured and this used to calculate the timing of the radiofrequency pulses automatically.
Minimising Eddy Currents
It is important that the units creating B0, B1 and the gradients do not interfere with each other via the creation of eddy currents. Eddy currents can be minimised by using a ceramic-ferrite permanent magnet for B0. The solenoid-like RF unit has also been designed to minimise eddy currents (see the above reference). However, in the present invention, eddy current effects can be minimised by exploiting the motional relativity principle. Namely, the field B0 and field gradients, Gz (and Gx, see below) are preferably kept constant in time and the object to be imaged is moved instead. The obvious exception is the radiofrequency unit which must be switched to create time-varying radiofrequency pulses.
The on-line imager uses various combinations of these basic hardware units to create images of the moving object using one or more of the on-line imaging protocols listed and described below. Which combination of units is used, and their arrangement along the conveyor, will be determined by the to choice of on-line imaging protocol which, in turn, will be determined in part by the nature and velocity of the object: to be imaged and the information required.
Because the NMR signal is acquired from an object moving with constant velocity v in a linear field gradient Gz oriented parallel to v, the conventional NMR pulse sequences will not, in general, succeed in giving image projections of the object. For example, Fourier transformation of the FID obtained with a 90xc2x0 pulse on the moving object in the field gradient will not give an image projection of the object. Nor, in general, will spin echoes or stimulated echoes be observed with conventional pulse sequences such as the Hahn echo (96-xcfx84-180-acquired), the CPMG or Stimulated echo sequences. (These pulse sequences are described in standard textbooks on MRI such as P. T. Callaghan, Principles of NMR microscopy, Oxford Science Publications, Oxford, 1991). Moreover, conventional Fourier transformation of the echoes would not give a image projection of the object. Special acquisition conditions and methods for extracting the image are therefore now described.
a) Creating a One-dimensional Image Projection from an MMFID
The digitised MMFID, acquired in quadrature, must first be transformed in the time domain by multiplication with the factor exp{xe2x88x92ixcex3G.vt2/2}, where t is the acquisition time after the 90xc2x0 pulse. This removes a phase factor created by motion with velocity v in a linear magnetic field gradient Gz The resulting transformed FID is then corrected for zero-order phase imbalance by adjusting the phase so as to give a zero first point in the imaginary part of the FID. A full echo is then formed by reflecting the FID using its complex conjugate. Finally the echo is Fourier transformed to obtain an image projection. A fast computer algorithm for achieving this is described in Appendix 2, together with simulated transforms.
b) Methods for Creating Image Contrast Using Motionally Modified Spin Echoes (MMSE""s) and Motionally Modified Free Induction Decays (MMFID""s)
In many applications, such as the on-line detection of bruises in fruit, it is desirable that images are acquired with contrast (or intensity distribution) weighted by one or more of several NMR parameters such as the relaxation times T2*, T2, T1, or the self-diffusion coefficient, D, or the flow velocity (in the case of fluids). The following protocols describe how motionally modified FID""s (MMFID""s) and/or motionally modified spin echoes (MMSE""s) can be used to create images weighted by each of these parameters:
T2 Contrast in Images Created Using Motionally Modified Spin Echoes
T2 contrast in images can be created using spin echoes created by hard 180xc2x0 pulses with a modified Hahn echo or CPMG (Carr Purcell Meiboom Gill) pulse sequence applied to the moving object in the linear field gradient Gz. As shown in Appendix 3, for a spin echo to be observed, the pulse spacing xcfx84 must be an integral multiple of the acquisition time AQ. Moreover, an image projection can only be extracted from the spin echo if it is first transformed to remove a motional phase factor. This is described in Appendix 4. A representative pulse sequence is shown in FIG. 6. T2 contrast is important because different parts of the object are often associated with different values of T2 and this will be seen in the image projection. This can be exploited in many ways. For example, frozen parts of foods have a much shorter T2 than unfrozen regions so that frozen and unfrozen parts of a food can be distinguished in the image.
T1 Contrast Using Motionally Modified Spin Echoes
There are several possible ways of introducing T1 contrast into the images. The simplest is to use the inversion recovery sequence whereby the polarized magnetisation, M(0) is inverted by a hard 180xc2x0 pulse and allowed to recover for a fixed time delay t1, adjusted for each application. After the time delay t1 an MMFID is created by a hard 90xc2x0 pulse generated by the RF unit in the presence of the linear gradient created by the Gz unit. The pulse sequence is illustrated in FIG. 7. If T1 is sufficiently long, the initial hard 180xc2x0 pulse can be eliminated by setting the polarizer unit to give B0 antiparallel to v. An alternative pulse sequence which has the advantage of permitting two images to be obtained and compared, one with T1 weighting, the other without involves the stimulated echo pulse sequence. A representative pulse sequence is shown in FIG. 8. Note that a motionally modified stimulated echo is only formed if T=nAQ, and that, because only longitudinal magnetisation exists between the second and third 90xc2x0 pulse (during the time t) the gradient Gz can be turned off between them. This has the advantage of permitting the use of two Gz units and two RF units rather than one very long unit.
Diffusion Contrast by Combining Motional Echoes with Motion-generated Pulsed field Gradients
Diffusion contrast can be created, explicitly, using two or more Gx units. Each Gx unit creates a constant, localised, non-uniform gradient Gx across the solenoid axis and transverse to the B0 field. Motion of the object through these static Gx gradients is equivalent to imposing time-varying (pulsed) field gradients. Suitable pulse sequences are shown in FIGS. 9 and 10. The first is based on the motionally modified Hahn echo pulse sequence; the second on a motionally modified stimulated-echo pulse sequence.
T1 (Low Field) Contrast using MMFID""s
With reference to FIG. 11, between the polarizing unit 3 and the imaging unit 4 the sample 2 contains longitudinal magnetisation relaxing in the earth""s magnetic field. By varying the time spent between the polarizer and detector, various amounts of T1(low field) relaxation can be introduced. A suitable arrangement is shown in FIG. 11.
c) On-line Detection of Defects and Changes in a Moving Object by Difference Imaging Using Motional Echoes
In some applications the acquisition of a single, parameter-weighted image is insufficient to identify a defect, bruise, foreign body etc. In such cases it may be necessary to take the difference between an image (called the first image) acquired with a contrast weighting which is insensitive to the presence of the defect and a second image weighted with a parameter (such as T2*, T2, T1, or D) which is changed by the presence of the defect. Taking suitably processed differences between the first and second images will then highlight the presence, position and extent of the defect.
The first image can be calculated from the MMFID generated by the first 90xc2x0 pulse in the pulse sequences listed above. The second image can be calculated from one (or more) of the motionally modified spin echoes generated subsequently.
The difference imaging protocol may, if necessary, be combined with image registration software and/or signal processing software, such as Wiener filtering. All such operations are performed in the least possible time using a fast computer, such as a 233 MHz PC.
d) 2 or 3 Dimensional On-line Imaging Protocols
There are several strategies for this:
Methods Based on Motionally Modified Spin Echoes
It is possible to undertake 3-dimensional imaging by making use of the Gxcfx86units to acquire MMSE""s in a uniform, steady magnetic field gradient Gxcfx86oriented transverse to B0 and to the direction of motion v. 3-dimensional imaging uses the back-projection technique, whereby the angular orientation xcfx86 of the field Gxcfx86 relative to the vertical (xcfx86=0) is incremented in equal angular steps between successive motional echoes (or Gxcfx86 units). A representative 3-dimensional pulse sequence is shown in FIG. 12. The image resolution in the transverse (x-y, or r,xcfx86 plane) is determined by the number of increments in the angle xcfx86 and hence by the number of Gxcfx86 units. The success of the protocol requires a long T2 so that sufficient spin echoes can be acquired, and the pulse spacing T must be an integral multiple of AQ.
Methods Based on the Motional Equivalence Principle
The Echo Planar Imaging (EPI) protocol uses rapidly switched gradients to generate an image of a non-translating object. Instead of a stationary object and rapidly switched gradients it is possible to create an equivalent pulse sequence by moving the object with linear velocity through an array of steady gradients. A suitable arrangement for two-dimensional imaging in the (x-y) plane transverse to the object""s velocity is shown in FIG. 13. Because the gradients are steady there is no problem with eddy current interference. The transverse gradients are created by locating Gxcfx86 units along the B0 and RF units.
e) On-line Flow Imaging Using MMFID""s
MMFID""s can also be used to determine the velocity distribution in fluids undergoing steady flow down a tube. This is useful for on-line measurements of rheological properties of fluids. Each volume element in a fluid in steady flow has a constant velocity, and so behaves as a very small rigid body and contributes its own motional echo at a position (or time) which depends on its velocity. By analysing the MMFID it is therefore possible to extract the velocity distribution in the fluid. The analysis is presented in greater detail in Appendix 7 for the special case that a thin slice of fluid is initially excited and that this slice is only a single voxel wide. In practice a thin slice of fluid in a plane perpendicular to the flow can be excited using a soft, shaped slice-selective radiofrequency pulse (which can be created by the RF unit) in the presence of the Gz gradient created by the Gz unit.
f) On Line MRI Temperature Mapping
Conventional MRI temperature mapping on stationary objects exploits the temperature dependence of NMR parameters such as the initial magnetisation M(0), the longitudinal relaxation time T1 or the diffusion coefficient D. The on-line protocols presented above also provide M(0), T1 or D weighted profiles of the moving object. By calibrating the temperature dependence of the M(0), T1 or D weighting it is therefore possible to detect on-line temperature changes in the object non-invasively. This could be of benefit in on-line monitoring of chilling or heating processes.
The present invention has thus far been described with reference to specific embodiments. It will be understood, however, that a number of modifications can be made thereto. For example, although the preferred embodiment requires that an object is undergoing continuous, uniform translational motion, in fact the object can be undergoing any non-zero velocity or finite acceleration providing that its movement can be precisely characterised such that the effect of the changing velocity on the nuclear magnetic resonance signals can be predetermined. For example, modification of the factor exp{xe2x88x92ixcex3G.vt2/2} used in transforming the MMFID signal would be required according to the precise velocity or acceleration of the object.
This has particular significance where objects to be imaged are falling off the end of a conveyor belt, for example. In that situation, the motion of an object can be precisely characterised where it is undergoing continuous acceleration. Arrival of the object into the imaging system and/or determination of its translational motion can be determined by optical beams, as discussed earlier.
Similarly, although it is desirable, as in the preferred embodiments, for the magnetic field B0 to be spatially uniform and for the magnetic field gradient Gz to be linear, it will be understood that these conditions need not hold providing that both the field B0 and the field Gz can be precisely pre-characterised such that appropriate compensation can be made for non-linearities in the transformations for image processing. This enables possible use of non-symmetrical, but precisely characterised coils.